The use of catheters is sharply increasing in a number of diagnostic and therapeutics procedures, leading to less invasive diagnostics and therapy. It is often required to perform pressure measurement of internal tissues or fluids. It is always of the utmost importance to have the capability of reducing the size of the catheter and thus, the pressure sensor must be as small as possible. Other important features for a pressure sensor are fidelity and stability.
Among the first invasive pressure sensors used in combination with catheters were external and were connected to internal body pressure by way of fluid filled tubing as described in U.S. Pat. No. 4,576,181, Wallace et al., 1986. This method suffers however from its extreme sensitivity against tubing movement which strongly affect the fidelity of the pressure measurement. In addition, for the tubing to properly transmit the dynamic of the pressure, it must be relatively large which provides poor compatibility with actual catheter design.
This important issue has been resolved with catheter tip pressure sensors such as those taught by U.S. Pat. No. 4,274,423, Mizuno et al., 1981 and U.S. Pat. No 4,771,782., Millar et al., 1988. These sensors can be made to fit into fairly small catheters without suffering from the sensitivity to catheter or patient movement. On the other hand, their use is limited to areas where no electromagnetic noise is present, such as in presence of MRI or electrosurgery. Another drawback of those electrical sensors is their sensitivity to moisture drift that results from the change of the conductivity of the media surrounding the pressure chip.
Fiber optic sensors have the potential of resolving those problems. Some initial design such as the one described by Matsumoto et al., “The development of a fibre optic catheter tip pressure transducer”, Journal of Medical Engineering & Technology, Vol. 2, no. 5 (1978) were based on the variation of the light intensity induced by various mechanism. Those sensors are well known to be prone to fluctuations due to all sort of environmental effects.
In the past few years, there have been an increasing number of fiber optic sensors based on the use of a Fabry-Perot cavity as the sensing element. Fabry-Perot sensors can be configured in various ways, such that they can measure a variety of parameters. In addition, it is possible to make low cost, miniature Fabry-Perot sensors by way of micromachining techniques (MEMS).
A variety of techniques have been proposed for interrogating a Fabry-Perot cavity. U.S Pat. No. 5,392,117, U.S. patent Filing Ser. No. 10/829,980 and U.S. Pat. Filling No. 60/610,950 are good examples of such interrogation techniques that have the ability to accurately measure a Fabry-Perot cavity length, i.e. the distance separating the two mirrors. It is worth mentioning that those interrogation methods are extremely accurate and reliable and thus, the final accuracy and repeatability of a Fabry-Perot pressure sensors used in combination with such an interrogation technique will strongly depend on the quality of the sensor itself.
Fabry-Perot based pressure sensors are then considered as those having the best potential to suit the needs for catheter tip pressure measurement. U.S. Pat. No. 4,678,904, Saaski et al., 1987 teaches one method of producing a Fabry-Perot sensor that has some very interesting characteristics. Although it can be produced at a fairly low cost while achieving good reproducibility, attaching the chip on tip of an optical fiber still have to be considered. For instance, bonding the pressure chip using solder glass makes the process very expensive, raising the cost to a level not compatible with medical uses. The pressure chip could be bonded on tip of the optical fiber using a polymer, but the cost would remain high because the tip of the fiber optic under this circumstance needs to be processed to receive the chip. For the chip to be well attached to the fiber, the fiber needs to be enlarged to a size comparable to the pressure chip with the addition of a tube. Notwithstanding the additional cost, the use of a polymer nearby the Fabry-Perot cavity makes this design prone to moisture induced drifting due to adhesive swelling.
Improvement to the previous design has been made by putting into practice the concept of chip level packaging. Gander et al. “Embedded micromachined fiber-optic Fabry-Perot pressure sensors in aerodynamics applications”, IEEE Sensors Journal, Vol. 3, No. 1 (2003) teaches a method of producing a pressure chip that includes an opening for accommodating the optical fiber, thus eliminating the need for additional fiber optic processing. On the other hand, this design can in no circumstance provide any stable pressure measurement. Since the tip of the optical fiber constitutes the first mirror of the Fabry-Perot cavity, the pressure measurement becomes functions of any fiber optic unavoidable pistoning due to moisture, temperature and handling.
Although the paper presented by Tohyama et al., “A fiber-optic pressure microsensor for biomedical applications”, Sensors and actuators, A66 (1998) disclosed a method of producing an intensity based pressure sensor, it teaches the use of a small silicon funnel produced on the chip level for accommodating the optical fiber. Kim et al., “Micromachined Fabry-Perot cavity pressure transducer with optical fiber interconnects”, SPIE Vol. 2642, (1995) teaches a similar concept, but using instead a Fabry-Perot cavity as the optical element. One important drawback of this method remains its sensitivity to moisture drift. It is known that polymers are swelling in presence of humidity. Upon swelling, the polymer that fills the silicon funnel for holding the optical fiber in place will induce a bending force to the pressure chip that brings the mirrors in closer proximity and thus, inducing drift.
FIG. 1 shows a prior art construction of a Fabry-Perot sensor 10 for measuring pressure. A bi-directional fiber optic 9 guides the light signal 7 toward a Fabry-Perot pressure chip 21. The pressure chip 21 is made of from a glass substrate 1. One first partially reflective mirror 2 is deposited within a recessed cavity 5 performed on the top surface of the glass substrate 1. A second deformable mirror 3 is bonded or welded to the glass substrate 1. Both mirrors 2, 3, spaced by a distance given by the depth of the recessed cavity 5, constitutes a Fabry-Perot cavity 6. The second mirror 3 bows toward first mirror 2 as function of an applied pressure. FP cavity length 6 is then an unambiguous function of pressure. The deflection of the diaphragm as a function of the pressure is usually of the order of 2 nm/mmHg. To be integrated into a catheter, the pressure chip 21 needs to be packaged so that the light signal travelling into the optical fiber 9 is directed from the fiber to the Fabry-Perot cavity 6, and back to the optical fiber 9. Although the use of an optical lens 8 can be considered, this does not yet resolve the method for attaching the fiber 9 to the glass substrate so that no environmental parasitic effect will be detrimental to the pressure measurement.
FIG. 2 illustrates another prior art where the fiber optic 9 is brought in close proximity to the Fabry-Perot cavity 6, usually less than 200 microns, so that no optical system is needed to bring the light in and out of the Fabry-Perot cavity 6. The presence of the adhesive 11 directly on the back side 12 of the glass substrate 1 makes this design very sensitive to moisture induced drift. Also, the insertion of the optical fiber 9 into the tube 13 contributes increasing the production cost.
FIG. 3A shows another prior art design where the method for assembling the fiber optic 9 to the pressure chip 21 is achieved at the chip level. The production cost is thus fairly acceptable. By preferential etching of silicon with methods well know by those skilled in the art, it is possible to micromachine funnels 22 into silicon substrate. With a (100) oriented silicon substrate, funnels with angles of 54 degrees are achieved. To adequately bond the fiber optic 9 into the funnel 22, it is required to completely fill the funnel with an adhesive 23. Considering the amount of adhesive required to fill a funnel with such a wide opening, the moisture induced swelling of the adhesive 23 will induce a strong bending moment 24, as illustrated by FIG. 3B, sufficient to bend the pressure chip so that the diaphragm 25 significantly deflects toward the glass surface 26, resulting in a pressure drift. Because the funnel is wide open, the pulling force required to pull the pressure chip off the fiber is significantly reduced.
Thus one drawback of current Fabry-Perot pressure sensors is their sensitivity to moisture.
Another drawback of Fabry-Perot pressure sensors is their sensitivity to temperature. For the sensors to find practical uses in medical fields, it is of importance to provide a method for compensating thermal shift that occurs as a result of the differences in the coefficient of thermal expansion of materials used to build the sensor.
Another drawback of the actual sensors is the variation in their sensitivity when immersed into a water based solution, phenomenon usually erroneously associated to moisture drift.